Magnetic resonance imaging (MRI), also called nuclear magnetic resonance (NMR) imaging, is a non-destructive method for the analysis of materials and is used extensively in medical imaging. It is completely non-invasive and does not involve ionizing radiation. In very general terms, nuclear magnetic moments are excited at specific spin precession frequencies which are proportional to the local magnetic field. The radio frequency (RF) signals resulting from the precession of these spins are received using pickup coils. By manipulating the magnetic fields, an array of signals is provided representing different regions of the volume. These are combined to produce a volumetric image of the nuclear spin density of the body.
In MRI, a body is subjected to a constant magnetic field. Another magnetic field, in the form of electromagnetic (RF) pulses, is applied orthogonally to the constant magnetic field. The RF pulses have a particular frequency that is chosen to affect particular atoms (typically hydrogen) in the body. The RF pulses excite the atoms, increasing the energy state of the atoms. After the pulse, the atoms relax and release RF emissions, corresponding to the RF pulses, which are measured and processed into images for display.
When hydrogen nuclei relax, the frequency that they transmit is positively correlated with the strength of the magnetic field surrounding them. A magnetic field gradient along the z-axis, called the “slice select gradient,” is set up when the RF pulse is applied, and is shut off when the RF pulse is turned off. This gradient causes the hydrogen nuclei at the high end of the gradient (where the magnetic field is strong) to precess at a high frequency (e.g., 26 MHz), and those at the low end (weak field) to precess at a lower frequency (e.g., 24 MHz). When the RF pulse, of a single frequency, is applied, only those nuclei which precess at that frequency will be tilted, to later relax and emit a radio transmission (i.e., the nuclei “resonate” to that frequency). For example, if the magnetic gradient caused hydrogen nuclei to precess at rates from 24 MHz at the low end of the gradient to 26 MHz at the high end, and the gradient were set up such that the high end was located at the patient's head and the bottom part at the patient's feet, then a 24 MHz RF pulse would excite the hydrogen nuclei in a slice near the feet, and a 26 MHz pulse would excite them in a slice near the head. Thus a single “slice” along the z-axis is selected; only the protons in this slice are excited to a higher energy level, to later relax to a lower energy level and emit a radio transmission.
The second dimension of the image is extracted with the help of a phase encoding gradient. Immediately after the RF pulse ceases, all of the nuclei in the activated slice are in phase. Left to their own devices, these vectors would relax. In MRI, however, the phase encoding gradient (in the y-dimension) is briefly applied, in order to cause the magnetic vectors of nuclei along different portions of the gradient to have a different phase advance.
After the RF pulse, slice select gradient, and phase encoding gradient have been turned off, the MRI instrument sets up a third magnetic field gradient, along the x-axis, called the “frequency gradient” or “read-out gradient.” This gradient causes the relaxing protons to differentially precess, so that the nuclei near the low end of the gradient begin to precess at a faster rate, and those at the high end pick up even more speed. When these nuclei relax again, the fastest ones (those which were at the high end of the gradient) will emit the highest frequency of radio waves. The frequency gradient is applied only when the signal is measured.
The second and third dimensions of the image are extracted by means of Fourier analysis. The entire procedure must be repeated multiple times in order to form an image with a good signal-to-noise ratio.
Finally, in spin-echo imaging, there is the problem that the inhomogeneity of the main magnetic field induces variations in the rate of precession of nuclei. To fix this problem, a 180° RF pulse is inserted into the cycle, at a time point halfway between the 90° pulse and the measurement of the radio transmission signal given off by the relaxing nuclei.
When 90° and 180° RF pulses are used, as described above, the technique is called “spin-echo imaging,” and takes several minutes to create a single image. Other techniques which can be used in MRI include “gradient echo imaging,” in which spin-echo's 180° pulse is replaced by a reversal of magnetic field gradients, “echo planar imaging,” a variation of spin-echo (also called “fast spin-echo imaging”), and “spin-echo inversion recovery imaging.”
MRI Techniques
In spin-echo imaging, an image section is defined by two RF pulses that are turned on following each other with a specified time interval. An echo signal is acquired once the same time interval has elapsed after the second RF pulse. The time between the first RF pulse and the signal is called the echo time TE. The echo signal results only when the tissue is exposed to both RF pulses, as is the case for stationary tissue. Blood flowing through the section will see the first pulse but move out of the section before the second RF pulse is given. Thus blood will give no signal or lower signal and appear darker than the surrounding tissue on the image. The intensity for blood is a function of the time TE/2 between the two pulses, the section thickness and the angle at which the blood flows through the section. For example, when the time between the pulses is 12.5 msec and the section thickness is 5 mm, blood entering the section perpendicular and moving with a velocity of 40 cm/sec will move completely out of the section in the 12.5 msec between the RF pulses and give no signal. Slower moving blood will move out only partially and appear with low but non-zero signal. Blood moving at an angle though the section, will take a longer time to traverse the section and may yield a non-zero signal.
Unlike spin-echo images, blood appears bright in fast gradient echo images. In gradient echo images, a stream of RF pulses is turned on in rapid time intervals with repetition times TR of 20-50 msec. When stationary tissue in an image section is exposed to this stream of RF pulses an equilibrium state is reached and results in a signal that is medium low (e.g., lower than the signal in spin-echo imaging). Moving blood enters the image section and is not exposed to the stream of RF pulses in the way the surrounding stationary tissue is. Thus moving blood gives a high signal. As in spin-echo imaging, the intensity of the blood signal depends on the section thickness, the velocity of the blood, the angle at which the blood enters the section and on the sequence acquisition parameters TR, TE and flip angle.
Physics of MRI
The proton nuclei of the hydrogen atom (and other atoms) possess a small magnetic moment. When placed within a magnetic field, a torque will be exerted upon them, resulting in a slight energetic advantage of one orientation (parallel to the field) over another (the anti-parallel orientation). Over time, random atomic collisions and other perturbations allow the complete system to reach a magnetic and thermal equilibrium with an excess of protons aligned with the magnetic field. The combined alignment of all of these protons results in a net magnetic moment; a subject placed within a magnetic field thus becomes “magnetized.” In biological tissues, this magnetization is exceedingly small, and generally not observable. In addition to their magnetic moment, atomic nuclei possess angular momentum—a quantum property known as “spin.” Because of this angular momentum, rather than aligning simply with magnetic fields, the individual nuclei precess about it, much as a spinning top or gyroscope might, when placed in the earth's gravitational field. The precessional rate, or frequency, is characteristic of the atomic nucleus (e.g., protons) and is proportional to the strength of the magnetic field, a property crucial to the process of image formation. With the magnetic field strengths in use for today's typical MRI machines, the precessional frequency is between 5 MHz and 200 MHz.
As shown in FIG. 1, the hydrogen proton 100 possesses the quantum property of “spin” or angular momentum 102, and has a small magnetic dipole moment 104. When placed in a magnetic field 106, a torque is exerted on the particle, causing it to precess about the applied field 108. The precessional frequency of the protons is directly proportional to the magnetic field strength. Protons precess at about 43 MHz/Tesla.
FIG. 2 shows that the proton magnetization can be decomposed into the sum of a stationary (longitudinal) 202 and a rotating (transverse) component 204. Each proton nucleus within a magnetic field thus yields a tiny field 206 that rotates about that applied field. The rotating field 206 from individual nuclei is generally aligned at random with respect to other protons in the subject or sample. In macroscopic systems, the average rotating field will effectively be zero, since that arising from any individual nucleus is canceled by another, oppositely oriented, neighbor.
In MR imaging, a second magnetic field is applied, which is orthogonal to the static field, and which rotates about the static field at the precessional frequency of the atomic nuclei. When the rotating field is present, the nuclei will precess about it, forcing the magnetization away from equilibrium, and causing the ensemble of protons to precess together, or in phase. The combined rotating magnetic moment thus produced by the ensemble of protons is observable as a time-varying electromagnetic signal. The second rotating magnetic field is applied at radio frequencies in the form of an RF pulse.
Two fundamental temporal parameters are used to describe the MR signal. The longitudinal relaxation rate, T1, is the rate at which nuclei, once placed in a magnetic field, exponentially approach thermal equilibrium, so that the magnetization (M) is described by the formula:M=Mo(1−e(−t/T1))  (1)where Mo is the equilibrium magnetization. In biological tissues, the proton T1 is quite long: from tens of milliseconds to seconds. Differences in the T1's of tissues are one of the primary bases of contrast in clinical MRI. A second parameter time constant describes the rate at which the MR signal decays. Once an MR signal is formed, i.e., after an RF pulse, it fades quickly; small variations in the local magnetic field, e.g., those caused by neighboring magnetic nuclei, cause the protons to precess at slightly different rates and therefore to become out of phase with one another. Interactions among the magnetized protons, and motion in inhomogeneous fields, due for example to diffusion, also results in signal dephasing. The observed signal decay rate, (T2*) generally ranges from a few milliseconds to tens of milliseconds and, to a reasonable approximation, also follows first order kinetics. The MR signal, S(t), signal decays according to the formula:S(t)=So(e(−t/T2*))  (2)where So, is the signal strength immediately following the RF excitation pulse. The observed T2* decay is the net effect of all the dephasing terms:
                                          1                          T              ⁢                                                          ⁢              2              *                                =                                    1                              T                ⁢                                                                  ⁢                2                                      +                          1                              T                ⁢                                                                  ⁢                2                ⁢                m                                      +                          1                              T                ⁢                                                                  ⁢                2                ⁢                D                                      +                          other              ⁢                                                          ⁢              terms                                      ,                            (        3        )            where T2m represents the dephasing due to magnetic field inhomogeneities and T2D is the diffusion-related signal loss. Like T1, the T2 signal decay rates differ among body tissues. By waiting for a period TE following the RF excitation pulse, differences in the signal decay rate will become evident as differences in the MR signal intensity: tissues with longer T2's will have stronger signals than those with short T2's, whose signals decay more rapidly. Modifications to the pattern of RF excitation (the “pulse sequence”) can modulate the contributions of the various relaxation processes to the resulting MR signal. In particular a spin-echo pulse sequence can be used to nearly eliminate the T2m contribution, increasing the relative contributions of other terms, such as proton diffusion, to the image contrast.
In each MRI technique, various parameters such as T1, T2, and T2* can be manipulated to increase or decrease the weight or value of the parameter in the MR image. However, each technique is affected by certain harmful physical characteristics, affecting signal-to-noise ratio, SNR. Each technique benefits from increasing SNR.